Orthopaedic reconstruction and trauma repair often utilize metallic, polymeric, and biodegradable materials devices, as well as allograft and autograft tissues, for replacement, fixation and repair of patient bones and other tissues. Some such devices are implanted percutaneously and temporarily, including fixator pins, while other such devices are surgically implanted and intended for long-term repair, replacement, and/or reconstruction, such as bone plates, spinal and dental implants, and artificial joints.
There are a variety of metal implants that are secured within bone by allowing bone to grow into pores purposefully manufactured on the surface of the implants. Early bone implants formed of metal with metal particles sintered thereon to create metallic porous surfaces are described in U.S. Pat. No. 3,808,606 by Tronzo. Such metal particles are typically sintered onto the implants surface, although other techniques of particle bonding are known including brazing as disclosed in U.S. Pat. Nos. 4,813,965 and 4,938,409 by Roberts. Other implants use porous coatings made from ceramic materials and hydroxyapatite, to encourage bone attachment, as detailed in U.S. Pat. Nos. 4,146,936 and 4,746,532 by Aoyagi et al. and Suzuki et al., respectively.
A number of orthopaedic devices are designed to be fixated by impaction, such as intramedullary nails which are driven into the marrow of bones. Other orthopaedic implants, for example bone plates and spinal interbody cages, are typically fixated with threaded screws. Many conventional implant coatings cannot withstand the shear forces associated with such impaction or torque fixation techniques.
All surgeries carry inherent risk of infection. Once an implant becomes colonized by microorganisms and a microbial biofilm is formed, eradicating such infection by systemic antibiotics is extremely difficult. Often an infected implant must be surgically removed, referred to as revision surgery, and the surrounding tissue and bone are debrided and treated directly via antibiotic lavage. Sometimes a short-term “spacer” element is temporarily implanted to hold anatomical structures in place while the patient undergoes antibiotic therapy. Ultimately a new long-term device is implanted. The cost, morbidity and mortality associated with such surgical intervention can be significant.
Reducing the risk of infection, therefore, is highly desirable. Surgeons often treat implants directly and/or the implant surgical site directly with antibiotics at the time of surgery in an effort to prophylactically protect the device and kill microorganisms inadvertently introduced into the operative site during the surgical procedure. Trauma-induced bone fractures often involve wounds that penetrate the patient's skin, breaching the skin's protective barrier, creating “dirty” wounds, requiring external fixator pins to align and fixate bone fragments percutaneously for months following injury until such broken bones heal. In such circumstances, microorganisms introduced at the time of the fracture, as well as the patient's own microbial flora migrating down a percutaneous fixator pin tract, can infect the patient's own soft tissue and bone, causing osteomyelitis and/or other soft tissue and blood stream related infections.
Reservoirs within implantable devices have been provided for antibiotics and other bioactive agents such as described in U.S. Pat. No. 6,709,379 by Brandau et al., U.S. Patent Publication Nos. 2006/0093646 by Cima et al. and 2007/0016163 by Santini Jr. et al. Other types of infection resistant orthopaedic devices include the EBI X Fix Dynafix™ System with SC Bone Screws which employs silver as a biocide, manufactured by Electro-Biology Inc. of Parsippany, N.J., and OrthoGuard™ AB Antibacterial Sleeve containing the antibiotic gentamicin, from Smith & Nephew, Inc. of Memphis, Term. Silver ions absorbed into hydroxyapatite coatings on metal implants are disclosed in U.S. Patent Publication No. 2009/0198344 by Prentice et al.
Orthopaedic implants are often secured with bone cement such as PMMA (polymethylmethacrylate), which may contain one or more organic antibiotics, such as Zimmer Holding's Palacos® R+G bone cement containing gentamicin and Stryker Corporation's Simplex™ P containing tobramycin. Antibiotic bone cements are also disclosed by Kuhn et al. in U.S. Patent Publication Nos. 2006/0292199 and 2008/0213336. Kuhn et al. teach that most of the antibiotic should be released within the first 24 hours after implantation.
Antibiotics are organic substances produced by microorganisms that are antagonistic to the growth of other microorganisms. Today, many such antibiotics are synthetically derived. The antagonistic property of an antibiotic is somewhat microorganism-specific. Consequently, the biocidal effectiveness of any antibiotic is highly predicated upon selecting the appropriate antibiotic for the infecting pathogen. Kuhn et al. promote the use of gentamicin as a “broadly effective” antibiotic in the above-cited U.S. Patent Publications, yet microbes, such as E. coli, have demonstrated some resistance to gentamicin, as have other microorganisms to other antibiotics. Also, possible side effects of gentamicin include damage to kidneys and/or nerves, which can cause dizziness, numbness, muscle twitching or seizures. Moreover, gentamicin is not recommended for people having kidney disease, hearing loss or loss of balance due to ear problems, or having a neuromuscular disorder.
Heat generated during the porous coating process of metal implants, various implant sterilization processes, and the elevated polymerization temperatures of a thermoplastic material such as PMMA, each may have deleterious effects on antibiotics and other organic materials and organic bioactive agents. While U.S. Pat. Nos. 5,876,446 and 5,947,893 by Agrawal et al. teach the use of antibiotic-loaded biodegradable polymers as a coating means for porous implants, the process of imbibing the antibiotic polymer into the porous coating is achieved via the use of solvents in an aseptic process, since the coating will not withstand steam sterilization temperatures. In a work titled “Bone Cements”, ISBN 3-540-67207-9, pages 16, 27-28, 141-142 and 254-258 (Springer, 2000), Dr. Klaus-Dieter Kuhn carefully distinguishes between the in-vitro polymerizing temperature of PMMA at approximately 80° C., from its in-situ temperature of approximately 46° C., from a tissue damage perspective, in particular, protein coagulation. However, Dr. Kuhn does not address the potential degradation of antibiotic effectiveness of gentamicin exposed to this temperature range. By comparison, the present inventor recognizes that inorganic compounds, both antimicrobial and osteoinductive, can withstand temperatures in excess of 400° C. and should remain therapeutically effective.
A method of treating porous coated prostheses with a biodegradable coating incorporating bioactive agents is described in the aforementioned Agrawal et al. patents. In those patents, a pharmacologically active substance, such as an antibiotic and/or an osteoinductive protein, is incorporated into a biodegradable polymer or calcium phosphate carrier and then is impregnated into the pores of the tissue-mating surfaces of the prosthesis. The biodegradable carrier dissolves over a period of weeks or months, releasing a drug to produce a pharmacological response or an osteoinductive material, e.g. bone morphogenic protein or trypsin inhibitor, which encourages bone in-growth.
Catheters having an outer coating layer which releases antimicrobial agents are described in U.S. Patent Publication No. 2008/0172011 by Heroux et al. and in U.S. Pat. No. 7,354,605 by Trogolo et al., for example. Some of these devices and other medical products utilize a water-soluble glassy material containing silver such as described in U.S. Pat. No. 5,470,585 by Gilchrist and U.S. Pat. Nos. 6,692,532 and 7,531,005 by Healy et al. Similar water-soluble antimicrobial compounds such as BACTIFREE™ and IONPURE™ water-soluble glasses are available from Mo-Sci Corporation of Rolla, Mo., and Ishizuka Glass Co., Ltd. of Nagoya, Japan, respectively.
A variety of thermoplastic polymers, due to their intrinsic biocompatibility and desirable physical properties, have found wide use as implantable biomaterials. Commonly utilized polymers include PMMA, mentioned hereinabove as ‘bone cement’ for fixating orthopaedic implants, and PEEK (polyether ether ketone), a colourless organic polymer thermoplastic, utilized for various procedures including interbody spinal fusion. Other implantable polymers include polyurethane and silicone, which are often employed when bending, flexing, or soft tissue interface is a prerequisite of implant utility and function. Such polymeric implants can be either molded, extruded, or thermoformed into final device configurations. Alternatively, machined polymeric implants start from block or rod forms which have been molded or extruded into initial work piece forms and are then processed into desired configurations.
Such implants commonly need to be internally fixated, whether for bone or soft tissue fixation. When such polymeric materials are interfaced with bone, often cement, screws, pins and plates are utilized to assist in polymer implant fixation. Also, such polymeric implant surfaces can be prepared with texture, grooves, undercuts or other surface features, which provides opportunity for bone ingrowth to further assist implant fixation. Osteoconductive materials, such as tricalcium phosphate, hydroxyapatite and biphasic calcium phosphate can be added to polymer resin, as filler material, in order to improve bone/implant fixation; however the ceramic particles used as filler material are generally small, approximately 10-50 microns, and poorly water soluble. Consequently any bone attachment is interfacial in nature possessing only modest shear strength.
DiFusion Technologies announced in January 2009 that it completed a series of laboratory tests of a certain silver ion-based antimicrobial technology in a PEEK spinal implant. As of the filing date of the present inventor's priority patent application, Provisional Application No. 61/301,698, the “CLEANFUSE MOA” mode of action section of the website “www.difusiontech.com” discloses that silver zeolite particles are compounded into the PEEK polymer during manufacture of the raw material prior to implant fabrication. Naturally occurring sodium ions in the bloodstream iontophoretically exchange with silver ions from the silver zeolite particles. The controlled cationic release of the silver ions, a broad spectrum biocide, allegedly kills hundreds of different types of microorganisms. The zeolite particles themselves, however, are not water-soluble and remain embedded within the PEEK polymer matrix without dissolving after surgical implantation.
Synthetic biodegradable polymers are often used in medicine when tissue scaffolds and/or controlled drug delivery is desired. As scaffolds, biodegradable polymers can be designed to approximate the physical and structural properties of the host tissue it replaces, while providing a suitable surface for host cell attachment and proliferation. Biodegradation of such material eliminates the need for removal of the scaffold or other device post-surgery. During the tissue remodelling process, healthy host tissue forms on the scaffold as the scaffold slowly biodegrades, ultimately leaving only native host tissue in its place. This technique can be used to regenerate new tissue such as skin, ligaments, tendons and bone. One such example is the L-C Ligament® by Soft Tissue Regeneration, Inc. in Connecticut. A braided biodegradable polymeric scaffold of poly L-lactic acid (PLLA) is implanted as an artificial cruciate ligament and is remodelled by the host's ligamentous tissue as the biopolymer biodegrades.
As drug delivery vehicles, biodegradable polymers act as a ‘time release’ capsule and reservoir for the bioactive agent impregnated within the polymer matrix. Drug eluting stents are a good example of the utility of such polymers in medicine.
Some of the more common synthetic biodegradable polymers include PGA (polyglycolide), PLA (polylactide), and PCA (polycaprolactone). Such polymers are synthesized by open ring polymerization and biodegrade because the polymer is hydrolytically unstable, i.e. water tends to open up and breakdown the polymer's structure. This action along with phagocytosis tends to eliminate the polymer from the patient's body. When used in drug delivery applications, biodegradable polymers are often combined in an effort to modulate the rate of hydrolysis. This in turn controls, to some extent, the elution kinetics of the bioactive agent.
Therefore the bioactive agent's elution kinetics, i.e. bioavailability, are based upon two factors: 1) the intrinsic solubility of the bioactive agent; and 2) hydrophobicity/hydrophilicity of the biodegradable polymer; for it is the hydrolytic property of the polymer that determines its: 1) biodegradation rate; and 2) the deteriorating rate of the material's mechanical, structural and physical properties.
It would be an error to assume that the erosion, that is, the removal or elimination, of the biodegradable polymer controls the elution kinetics of the bioactive agent, in other words, that the bioactive agent is still present and effective as long as the biodegradable polymer is present. This is most often not the case, since the solubility of the bioactive agent is generally far more soluble than the biodegradable polymer. Therefore, once moisture penetrates the polymer as part of the hydrolysis/biodegradation process, the same moisture solubilises, that is, dissolves, the bioactive agent, causing it to elute from the polymer. While it can be said that biodegradable polymers do in fact slow down the bioactive agent elution process, it would be fallacious to assume that the biodegradation process of the polymer was equivalent to the bioavailability of the bioactive agent within the polymer.
Biodegradable polymers therefore can retard, and to some extent control, the bioavailability of the bioactive agent. However in most clinical situations the bioactive agent elutes at a considerably faster rate that the biopolymer biodegrades. Typically, the bioactive agent is exhausted well before the polymer biodegrades.
Use of a bioactive material in a bone-implant prosthesis is described in U.S. Patent Publication No. 2003/0171820 by Wilshaw et al., for example. One preferred bioactive material is disclosed as “Bioglass® 45S5”, a calcium phospho-silicate bioactive glass or glass-ceramic material, also referred to by others and herein as “45S5 Bioglass®” bioactive glass. Further examples of implants with bioactive glass are provided in U.S. Patent Publication Nos. 2008/0038534 and 2009/0208428 by Zenati et al. and Hill et al., respectively.
Mixtures including 45S5 Bioglass® material are promoted by Novobone Products, LLC of Alachua, Fla. as PerioGlas®, a synthetic absorbable, allegedly osteostimulative bone graft substitute which is mixed with sterile water, saline, blood or marrow from a patient, or with autogenous or allograft bone particles to create a paste that is applied to dental intraosseous, oral or cranio-/maxillofacial bony defects. Similarly, Nanotherapeutics, Inc., also of Alachua, Fla., provides ORIGEN DBM with bioactive glass as a malleable bone-void filler. The filler comprises human demineralised bone matrix (“DBM”) and synthetic calcium phospho-silicate particles, both coated with porcine gelatin.
Other uses of 45S5 Bioglass® material include coatings for sutures with and without silver, as described by Pratten et al. in “In vitro attachment of Staphylococcus epidermis to surgical sutures with and without Ag-containing bioactive glass coating”, J. Biomater Appl. vol. 19, pages 47-57 (2004). Glass-ceramic scaffolds formed from sintered particles of 45S5 Bioglass® which encourage bone tissue in-growth are explained by Chen et al. in “45S5 Bioglass®-derived glass-ceramic scaffolds for bone tissue engineering”, Biomaterials vol. 27, pages 2414-2425 (2006).
It is therefore desirable to have improved techniques for promoting in-growth of bone or other tissue into an implant, into implant coatings, and into bone cement, preferably while also reducing the inherent risks of surgical site and implant infection.